Effective scaffolding is crucial to the success of all tissue-engineering applications and ex vivo cell expansion applications. The design of effective scaffolds has recently been focused on incorporation of specific matrix chemistry, substrate surface configuration and three-dimensional macrostructure design. Polymer scaffolds must possess several key characteristics, including high porosity and surface area, structural strength, and specific three-dimensional shapes, to be useful for tissue engineering applications.
Developing polymeric scaffolds with high porosity, i.e. high surface to volume ratio to provide a large amount of surface for cell attachment has been one of the most active research topics. Several techniques have been established for processing polymers into a porous structure. Most of these methods are based on a class of biodegradable polymers, poly(lactic acid) (PLA), poly(glycolic acid) (PGA) and their polymers (PLGA). Particulate leaching is the first method that has been employed for the fabrication of biodegradable porous foams. This method, however, has less control of the microarchitecture of the pore structure and uniform porosity. An obvious limitation is the difficulties of scaling up of this fabrication technique (Mikos, et al. 1993; Ma, et al. 1998).
Recently, textile technologies are used to fabricate biodegradable woven or nonwoven fabrics as tissue engineering scaffolds (Ma, et al. 1995). Fibers provide a large surface area to volume ratio and therefore are desirable as scaffold materials. The first studied fabric scaffold is a nonwoven mesh made of PGA sutures. Nonwoven PGA fibrous matrix is prepared by entangling fibers or filaments to form an isotropic 3-D matrix structure, leaving a space with a high void volume and a typical porosity in the range of 80-90%. These fibrous matrix lacks of structural stability necessary for the cell culture use. Therefore, several fiber-bonding techniques have been developed to prepare the interconnected fiber networks with different shapes as tissue engineering scaffolds (Thomson, et al. 2000).
Nonwoven fabrics design, compared with biodegradable foams formed by particulate leaching, offers a better control over the scaffold porosity and the fabrication process is more reproducible. These nonwoven mesh scaffolds have achieved good success in several tissue engineering applications, including urinary bladder (Oberpenning, et al. 1999), vascular graft (Niklason, et al. 1999), Trileaflet Heart Valves (Hoerstrup, et al. 2000), cardiac graft (Li, et al. 2000), skeletal muscle (Saxena, et al. 1999), cartilage (Naumann, et al. 1998), etc. Nevertheless, the current available scaffold designs using polymer fibers (mostly non-woven mesh) still pose several limitations.
Firstly, the surface of the fibers used to fabricate scaffolds or matrixes lacks of functional ligands required for cell attachment, proliferation and function. PGA fiber surfaces are not the natural substrate for cell attachment and growth. In almost all the studies mentioned above, the non-woven meshes have been coated by another biodegradable polymer as a binder (e.g. poly-4-hydrobutyrate, PHB) or treated by partial alkali hydrolysis to modify the adsorption of serum proteins onto the surface-hydrolyzed fibers to improve cell attachment and seeding density (Gao, et al. 1998). This process would affect the degradation kinetics of the biodegradable fibers, and is also much less controllable. Moreover, the modified surface adsorbed with serum proteins has no specificity to cell types. Similar approach is taken for non-degradable fibrous matrix. Polyethylene terephtahlate (PET) fibers are partially hydrolyzed and to create enough functionalities on fiber surface to enhance the attachment of the extracellular proteins and therefore improve cell adhesion (Ma, et al. 1999). This patent provides methods to conjugate bioactive signal proteins to the surface of biodegradable fibers and non-degradable fibers.
Secondly, polymer materials used to process biodegradable fibrous scaffolds have been limited to PGA although different bonding materials have been used to stabilize the scaffolds, mostly PLA or PHB. The degradation products of PLA, PGA and PLGA are glycolic acid and lactic acid. They would create an acidic microenvironment at the cell-scaffold interface. Low pH microenvironment is known to be detrimental to maturation of many types of cells and tissue development. Shum-Tim et al. have engineered an ovine pulmonary valve leaflet and the pulmonary arteries from autologous cells using nonwoven PGA mesh (Shum-Tim, et al. 1999). Use of this cell-polymer construct in the systemic circulation resulted in aneurysm formation. This is possibly due to the acidic degradation products or lacking the structural integrity throughout the remodeling process. New biodegradable materials suitable for fiber processing are in great demand to overcome this limitation. This patent also provides a serious of new biodegradable materials that could be processed into fibers and amendable to surface conjugation.
Lastly, nonwoven fabric designs lack of the control of scaffold microarchitecture. Obtaining a uniform porosity is not possible. In addition, nonwoven fabric scaffolds generally have weak mechanical structures. Certain bonding or backing materials are needed to ensure the structural stability. Examples of structural re-enforcing techniques include polypropylene fiber backing for PET meshes (Wang, et al. 1992), solution coating or spray coating of a PLA or PLGA layer (Mikos et al. 1993; Mooney, et al. 1996), sewing with Dexon suture (Niklason et al. 1999), and polyglactin suture (Oberpenning et al. 1999) for PGA meshes. This patent provides methods using textile technologies to provide scaffolds with coherent and ordered structures. Polymer fibers are woven or knitted to form three-dimensional scaffolds with different designed pattern to obtain various degrees of porosity (Wintermantel, et al. 1996), microtopology of the cell culture environment and microdistribution of the functional ligands using surface modified fibers.
This patent describes methods of preparing biofunctional fibers based on non-degradable fibers and biodegradable fibers, describes a serious of new biodegradable materials that could be processed into fibers and amendable to surface conjugation, describes methods of preparing fibrous scaffolds by surface biofunctionalization or using biofunctionalized fibers. These technologies will find wide applications in tissue-engineering and bioprocessing fields. Two specific examples are illustrated below to demonstrate the advantages of this scaffolding technology—stern cell expansion for nondegradable fibrous scaffolds, and vascular graft engineering for the biodegradable scaffolds.
1. Current Stem Cell Expansion Methodologies
A technology for efficient and practical ex vivo expansion of hematopoietic stem cells and progenitor cells would find wide applications in stem cell transplantation and somatic gene therapy. For detailed clinical applications of the expanded haemopoietic progenitor cells, see reference (Alcorn, et al. 1996). Current methodologies for ex vivo stem cell expansion are still far from optimal in achieving high expansion rate and maintaining pluripotency.
The goal of ex vivo expansion is to induce cell division and proliferation of stem cells while maintaining their primary functional phenotypes, namely, their ability to engraft and sustain long-term hematopoiesis. Over the past few years, techniques have become available that allow the extensive proliferation of haemopoietic progenitor cells in ex vivo culture systems. One method of stem cell expansion utilizes an adherent monolayer of stromal cell, which supports the viability of stem cells and early progenitor cells (Dexter, et al. 1977). Briefly, in the first few weeks of culture, a complex adherent layer of stromal cells is laid down. This stromal layer comprises fibroblasts, macrophages, adipocytes, endothelial cells and reticular cells. Hematopoesis can be maintained for months in a long-term bone marrow culture and it is thought that direct adhesive interactions between the hematopoietic cells and various elements of the stroma are crucial to the regulation of primitive hematopoietic cells. This suggests that the complex stromal layer can, to some extent, successfully mimic the unique microenvironment present in the bone marrow. The major advantage of these stromal-based culture systems is their ability to expand the numbers of primitive hematopoietic cells.
Although stromal layer may provide a suitable substrate for hematopoietic cell immobilization and culture, it has a number of limitations. The stromal layer is fragile. Therefore, it requires a rigid substrate on which the layers of stromal cells should be grown in order to maintain the integrity of the stroma. Moreover, cells grown on stroma only have a limited culturing lifetime of about six to eight weeks due to death of the stromal cells. More importantly, the use of stroma for a clinical ex vivo application poses a considerable logistic problem. In most cases, the stromal cells are obtained from the patient to avoid the immuno-rejection. The need to first collect and then grow a layer of the patient's stromal cells before they can be used to culture the hematopoielic cells adds to the time, cost, and complexity of the production of the autologous HPC cells. Moreover the stromal layers are much less defined. It introduces an additional highly variable factor into the culture system. This renders the controlled culturing difficult if reproducible stromal cultures of predictable characteristics are to be obtained. Allogeneic source of stroma, although feasible, is unreliable. The fact that a primary allogeneic stroma has to be irradiated suffers, as any donor-derived tissues would, the potential risks of infection. The quantity to which primary stromal cells can be expanded is limited. Immortalized human stromal cell lines are potentially unlimited in quantity (Roecklein, et al. 1995). However, no allogeneic stromal support is currently available that is suitable for clinical use yet (von Kalle, et al. 1998).
For these reasons, ex vivo culture of HSCs in suspension without stroma a has been actively pursued in recent years. The most widely used method for ex vivo expansion has been a relatively simple liquid suspension culture system supplemented with a combination of a range of cytokines (Hoffman, et al. 1995). The development of HSC in vivo is thought to be regulated, at least in part, by interactions of cytokine receptor signals. Various combinations of cytokines have therefore been. studied to obtain the optimal culture conditions for HSC expansion. In particular, stem cell factor (SCF) and Flk-2/Flt-3 ligand (FL) have been used as key cytokines for HSC expansion, because c-Kit and Flk-2/Flt-3, tyrosine kinase receptors for SCF and FL, respectively, have been shown to transduce signals crucial for HSC development. Thrombopoietin (TPO), a ligand for c-Mpl, originally identified as a primary regulator for megakaryopoiesis, has also been shown to stimulate the expansion of primitive hematopoietic cells. A recent study showed that a combination of SCF, FL, TPO, and a complex of IL-6 and soluble IL-6 receptor (IL-6/sIL-6R), was able to induce a significant ex vivo expansion of human hematopoietic stem cells for 7 days. The expanded cells were capable of repopulating in NOD/SCID mice, leading to successful bone marrow engraftment in the recipient animals as measured by considerable numbers of human CD45+ cells 10-12 weeks after transplantation (Ueda, et al. 2000). Simplicity is a major advantage of the cytokine-supplemented suspension culture. In a typical process, CD34+ cells are suspended in culture medium and incubated in an appropriate vessel (tissue culture flasks (Brugger, et al. 1995) or gas-permeable culture bags (Alcorn, et al. 1996; Mellado-Damas, et al. 1999)) for between eight to twelve days. The culture cells can then be harvested with ease and used as required. The medium is preferably serum-free, although a number of studies have used serum-supplemented medium. Serum-free culture allows the researcher to develop a chemically defined medium with known amount of cytokines, therefore the cell expansion process is more controlled and reproducible, and easy to scale up. More importantly, the use of serum free conditions is highly recommended for cell therapy protocols such as employing HPC-derived dendritic cells (DC) and T cells, whose exposure to exogenous antigens can be limited to a minimal level.
While the general protocols for suspension culture are similar, there are a variety of different cytokine recipes developed by various groups. The cytokines most commonly used include a combination of SCF, Flt-3 Ligand, TPO, G-CSF, GM-CSF, IL-3, IL-6, and erythropoietin (Epo). Several recent studies have suggested that SCF, Flt-3 ligand, TPO, and IL-3 might play key roles in the early human hematopoiesis. The combination of these cytokines (especially Flt-3 ligand and TPO) significantly enhanced the amplification of primitive HSCs (Petzer, et al. 1996; Petzer, et al. 1996; Piacibello, et al. 1997; Yagi, et al. 1999). The degree of ex vivo expansion is normally assessed by calculating the fold-increase in total numbers of cells, committed progenitors, CD34+ cells, and LTBMC-IC with respect to the input cells. Routinely, extensive expansion of cell numbers is obtained. Depending on the duration of culture, this can vary from a 30-fold increase in cell numbers from an eight-day culture, up to over 1000-fold increases with longer periods of 14 to 21 days. Similarly, numbers of committed progenitor cells also increase, for example, 41-fold following an eight-day culture, up to 190-fold from a 14-day culture. By repeated feeding of cultures, cell numbers can continue to increase for up to 21 days.
Generally speaking, no stromal influence is incorporated into the suspension culture system, although various combinations of cytokines are utilized to provide the proliferation and differentiation signals that stroma is thought to deliver. The addition of cytokines is thought to compensate for the absence of stroma-associated support. This represents a major disadvantage when one considers that, in vivo, blood cell production is regulated at a local level by interactions of hematopoietic stem cells with a variety of cell-bound and secreted factors produced by adjacent bone marrow stromal cells. It is unlikely that the cytokine combination currently in use will be adequate substitutes for stroma.
Another limitation of the serum-free suspension culture is the low expansion of the true stem cells, which is measured by long-term-culture-initiating cell (LTC-IC) assay. There is little evidence of significant LTC-IC proliferation, with, at best, maintenance of LTC-IC numbers over the culture period under these conditions. This is probably related to the fact that the current system lacks the unique regulatory microenvironment of bone marrow stroma. Nevertheless, a recent study showed that using a much higher concentration (30-fold higher) of cytokines than for maximal amplification of colony-forming cells, a 60-fold expansion of LTC-ICs from primitive cells has been achieved (Zandstra, et al. 1997). However, other studies have suggested the induction of differentiation of murine stem cells and thus loss of their repopulating ability when high concentration of IL-1, IL-3 and IL-6 are used for the ex vivo expansion (Jonsson, et al. 1997). Down regulation of surface IL-3 receptor in response to the high concentration of soluble IL-3 may have played a role. Immobilized HGFs may alleviate this problem by only providing high concentration of growth factors at the “reaction site”.
Recent insights into hematopoietic stem cell biology have demonstrated that the three-dimensional architecture of the culture environment may influence the maintenance of stem cell pluripotency in vitro. Several studies employing three-dimensional devices made of synthetic polymers support the hypothesis that physical topography of bone marrow microenvironments plays an important role in maintaining hematopoietic stem cell viability and pluripotency (Naughton, et al. 1989; Naughton, et al. 1990). These studies show that a 3-D microenvironment supports HPC survival, proliferation and multilineage differentiation. Naughton and Naughton have developed a three-dimensional cell culture apparatus for HSC expansion, in which a stromal support matrix is pre-estabilished and grown on the polymeric mesh surface (Naughton, et al. 1992). An interesting study by Rosenzweig et al. indicates that culturing hematopoietic progenitor cells (HPCs) in a three-dimensional tantalium-coated porous biomaterial structure enhances HPC survival, and preserves primitive CD34+ CD3831 cells, even without using hematopoietic growth factors as compared with standard culture techniques. This culture technique improves retroviral transduction of CD34+ cells and LTC-ICs without loss of multipotency (Rosenzweig, et al. 1997).
In summary, other than defining the source of HSCs and developing methods to obtain a purer CD34+ cell source, optimizing the ex vivo culture methodology represents the major challenge for HSC expansion. Considering the various aspects of ex vivo culture of HSCs, we hypothesize that a successful new generation of HSC culture system should include the following key features: (1) a three-dimensional culture device that mimic the microenvironment in the bone marrow stroma, (2) matrix-bound cytokines (including SCF, Flt-3 ligand, TPO, etc.) that mimic the in vivo configuration where these crucial cytokines interact with HSCs in vivo in early hematopolesis, (3) a bioreactor system that is easy to scale up to obtain a clinically acceptable expanded stem cell population.
2 Tissue Engineering of Small Diameter Vascular Grafts
Surgical treatment of vascular disease is now a common medical procedure. However, to date, the use of synthetic polymeric materials is limited to grafts larger than 5-6 mm due to the frequency of occlusion observed with synthetic vessels of smaller diameters. Consequently, significant efforts in the past 15 years have been focused on the development of a small-diameter blood vessel equivalent using tissue-engineering approach. The seeding of synthetic grafts with endothelial cells has been investigated as a means to increase patency, but has been limited by the challenges associated with maintaining effective surface coverage. As an alternative to the use of synthetic materials, two approaches have been taken to create a blood vessel using cell and matrix components. One approach is to create an acellular graft constructed of a material, such as collagen, that would provide the required mechanical properties on implant but would also facilitate remodeling and infiltration of host cells into a cellular vessel (Sullivan, et al. 2000). In this approach, the acellular matrix allografts or xenografts often times require a crosslinking process to provide the requisite mechanical characteristics, and the potential inflammatory response to the acellular grafts still persists. Another approach has gain great attention recently, uses techniques to create a cellular vessel through culture of smooth muscle cells within a biodegradable fibrous matrix and lining the lumen with endothelial cells (Niklason, et al. 1997; Shinoka, et al. 1998; Zund, et al. 1998; Niklason et al. 1999; Shum-Tim et al. 1999).
Weinberg C B and Bell E have first demonstrated in vitro development of a model blood vessel in a porous collagen scaffold in 1986. The remodeled blood vessel has three layers corresponding to an intima, media, and adventitia (Weinberg, et al. 1986). A confluent layer of endothelial cells was grown in culture onto the lumen of a tubular collagen construct consisting of an outer layer of fibroblasts and a middle layer of smooth muscle cells. An external Dacron mesh was used to provide additional mechanical support. However, elastin, the principal arterial-tissue-matrix protein besides collagen, was not present in the model. Matsuda T and Miwa H also created a hybrid construct using a polyureathane scaffold seeded with smooth muscle and endothelial cells (Matsuda, et al. 1995). This construct was shown to remodel in vivo successsfully in a canine model for up to 1 year. It is worth noting that in both of these two designs, a nondegradable polymer support was used to reinforce the strength of the cellular layers.
The state-of-art scaffolding technology in tissue engineering of blood vessel is to employ synthetic nonwoven biodegradable fibrous meshes. Using a partially hydrolyzed PGA nonwoven fabric scaffold, Niklason L E et al. have cultured bovine vessels under pulsatile media flow conditions (Niklason et al. 1999). In this study, vascular biopsy derived aortic smooth muscle cells have been seeded in the scaffold and cultured for 8 weeks, before seeding the endothelial cells in the luminal surface. Pulsatile radical stress is applied to the vessels at 165 beats per minute and 5% radical distention. The remodeled vessels have rupture strengths greater than 2000 mmHg and suture retention strengths of up to 90 grams, and exhibit the beginnings of vascular contractile responses. These engineered arteries have been implanted in miniature swine, and remain patent for up to 3 weeks postimplantation. However, these engineered vessels are also notably lacking in elastin content. In another in vivo blood vessel engineering model, Shum-Tim D et al. have reported a tissue engineered ovine pulmonary artery from autologous cells cultured in a PGA fibrous scaffold (nonwoven mesh) (Shum-Tim et al. 1999). Polyhydroxyalkanoate (PHA) layers have been used to provide the temporary mechanical characteristics of the tubular scaffold as the cells lay down their own extracellular matrix on the PGA surface, which ultimately takes over the structural integrity and biomechanical profile of the engineered tissue. Ovine carotid arteries are harvested, expanded in vitro, and seeded onto 7-mm diameter PHA-PGA tubular scaffolds. The autologous cell-polymer vascular constructs have been used to replace 3-4 cm abdominal aortic segments in lambs. All tissue-engineered grafts remain patent for up to 5 months, and no aneurysms developed by the time of sacrifice. The mechanical strain-stress curve of the TE aorta approaches that of the native vessel. In both studies, scaffolds have been used without any cell adhesive molecules on the surface. A bioadhesive surface would obviously increase the cell seeding efficiency and shorten the time needed for in vitro modeling. This has been difficult to achieve using the current available polymeric materials.
Another key challenge in developing a tissue-engineered blood vessel is to create a construct with the required mechanical properties. Several studies have demonstrated that optimizing the in vitro culture conditions would increase the burst strength of the engineered blood vessel. A few factors that would significantly affect the mechanical characteristics of the remodeled blood vessels include media flow (Ziegler, et al. 1995), ascorbic acid supplement (L'Heureux, et al. 1998)), glycation of the media equivalents (Girton, et al. 1999; Girton, et al. 2000), and particularly, applying pulsatile mechanical stimulus to the cellularized constructs (Niklason et al. 1999). This requires a scaffold with good mechanical strength, which nonwoven-mesh scaffold lacks. As an alternative, additional biodegradable suture, coating or silicon tubing has been used to provide structural integrity and mechanical properties for these non-woven mesh scaffolds (Niklason et al. 1999; Oberpenning et al. 1999; Shum-Tim et al. 1999).
This patent provides biodegradable polymers with functional side chains for the conjugation of adhesion molecules, provides methods of preparing fibrous scaffolds based on biofunctional fibers derived from these polymers.